1. Field of the Invention
The invention relates generally to nanoparticles that are used to deliver drugs into the cytoplasm of cancer cells and, more specifically, to the rapid delivery of the drugs by nanoparticles that have an inner core containing an anticancer drug and a polymer that is soluble in cancer cells.
2. Background of the Art
Cancer is the second leading cause of death in the United States. Each year more than 1.2 million Americans are diagnosed with cancer, and less than half can survive five years. Annual medical costs for cancer treatment account for billions of dollars in the US alone. Chemotherapy, which uses chemical agents (anticancer drugs) to kill cancer cells, is one of the primary methods of cancer treatment. Unfortunately, these anticancer drugs have limited selectivity for cancer and are inherently toxic to both cancer and normal tissues. As a result, anticancer drugs can cause severe side effects and damage to healthy tissues. For example cisplatin is a well-known metal complex that exhibits high antitumor activity [Rosenberg et al., 1969; Takahara et al., 1995]. However, it has significant toxicity, in particular, acute as well as chronic nephrotoxicity [von Hoff et al., 1979; Pinzani et al., 1994]. Other common side effects of anticancer drugs include decrease in the number of white blood cells (increasing risk of infection), red blood cells (losing energy) and platelets (risk for bruising and bleeding) as well as nausea, vomiting, hair loss, etc. Furthermore, the high glomerular clearance of the anticancer drugs leads to an extremely short circulation period in the blood compartment [Siddik et al., 1987].
Most importantly, treatments in conventional dosage form of these drugs may lead to initial cancer regression, but soon the cancer becomes insensitive to the drugs, causing cancer progression and death. The primary reason for the treatment failure is cancer's intrinsic and acquired drug resistance [Pastan and Gottesman, 1991; Gottesman, 2002]. When a conventional drug dose is administered intravenously, the drug molecules distribute throughout the body and some drug molecules reach the cancer interstititium. Some are taken up by cancer cells via diffusion, transport and endocytosis. On the other hand, cancer cells have various mechanisms by which they become resistant to the drugs, such as loss of a cell surface receptor or transporter for a drug to slow down the drug influx, specific metabolism of a drug, alteration by mutation or drug detoxification to consume the drugs, and the like [Gottesman, 2002]. A major mechanism of multidrug resistance is an energy-dependent drug efflux transporter, the P-glycoprotein (P-gp) pump located in cell membrane [Gottesman, 2002]. P-gp pumps are very efficient in detecting and binding a large variety of hydrophobic drugs as they enter the plasma membrane. These pumps then transport the drugs out of the cells [Bogman et al., 2001; Gottesman, 2002]. As a consequence of the slowed drug entry but efficient drug removal by the P-gp pumps and the drug consumption by other forms of drug resistance, the effective drug concentration in cytoplasm is well below the cell-killing threshold, resulting in a limited therapeutic efficacy.
Thus, a continuing challenge in cancer treatment is to develop new methodologies that have great drug selectivity for cancer and overcome the cancer drug resistance to simultaneously enhance the therapeutic efficacy and reduce toxicity to healthy tissues.
It has been demonstrated that cancer-targeted drug delivery, which preferentially delivers drugs to cancer tissues, can substantially reduce drug toxicity and enhance the therapeutic efficacy. The cancer-targeting is achieved by passive accumulation through cancer's leaky blood capillaries Hobbs et al., 1998; Monsky et al. 1999; Maeda, 2001; Jain, 2001; Torchilin, 2001]. The pore cutoff size of cancer's blood capillaries was reported ranging between 380 and 780 nm [Hobbs et al., 1998; Yuan et al., 1995] or around 400 nm [Unezaki et al., 1996]. This leaky nature allows for easier extravasation of larger molecules or colloid particles to the cancer tissues. In addition, cancer also has much fewer lymphatic capillaries than healthy tissues, such that the lymphatic drainage of macromolecules from cancer tissues is inadequate. As a result of the hyperpermeability of cancer vasculature and the absence of lymphatic drainage, macromolecules or colloidal particles are passively trapped in cancer tissues. This is referred to as the “enhanced permeability and retention effect” (EPR) [Maeda et al., 2001; Lukyanov et al., 2002]. This does not happen as much in healthy tissues because the much tighter blood vessels openings (just several nm) [Seymour, 1992] are almost impermeable for macromolecules and colloid particles. Active cancer-targeting by receptor-mediated delivery has also been achieved such as folic acid-mediated delivery [Lu et al, 2002; Gosselin and Lee, 2002]. The resulting drug concentration in the tumor can be several to tens of times higher than those in healthy tissues [Seymour, 1992; Lukyanov et al., 2002].
Of the various approaches developed for targeted drug delivery, polymer nanoparticle technique has been attracting increasing attention since it offers suitable means to deliver drugs to tissues or cells [Labhasetwar et al., 1997; Kwon, 1998; Brigger et al., 2002; Hans and Lowman, 2002]. Nanoparticles are referred to as submicron colloidal particles. Due to the subcellular size, they can penetrate through fine capillaries, cross the fenestration into interstitial space, and are easily taken up by cells via endocytosis/phagocytosis. Furthermore, nanoparticles (less than 100 nm) with a hydrophilic surface, such as a poly(ethylene glycol) (PEG) layer, can evade the recognition and subsequent uptake by the reticuloendothelial systems (RES) and thus have a prolonged circulation in the blood compartment, which is needed for the passive accumulation in cancer tissues via EPR [Gref, et al., 1994; Bogdanov et al., 1997; Moghimi et al., 2001; Kaul and Amiji, 2002]. Certain types of nanoparticles were also found to be able to overcome multidrug resistance to some extent [Brigger et al., 2002]. Nanoparticles are also much stable than liposomes and thus preclude the breakage in the bloodstream. Drugs can be physically entrapped in the core and do not experience harsh reactions. Nanoparticles also have large surfaces that can be used to modify the surface properties such as attachment of targeting ligands for site specificity.
Nanoparticles with long-circulation-times, also called stealth nanoparticles, can be fabricated from micelles formed by self-assembly of amphiphilic copolymers [Kreuter, 1994; Kwon and Kataoka, 1995; Kwon, 1998; Kataoka et al., 2001]. Such nanoparticles have a core-shell structure. The hydrophobic inner core has a high drug-loading capacity. The tight hydrophilic shell (usually composed of PEG chains) prevents the interaction of the hydrophobic core from protein adsorption and cellular adhesion and thus protects the drug in the core from hydrolysis and enzymatic degradation. The PEG chains also prevent the recognition by the RES [Moghimi et al., 2001; Brigger et al., 2002]. Thus, these so-called ‘stealth’ properties of the PEG shell result in an increased blood circulation time of the nanoparticles and allow drugs to passively accumulate in tumor tissues by EPR effect [Moghimi et al., 2001; Brigger et al., 2002]. These nanoparticles have been used as anticancer drug carriers such as cisplatin [Yokoyama, et al., 1996; Bogdanov et al., 1997].
The prior art has several disadvantages or drawbacks. First, the premature burst release of drugs in bloodstream is a general problem of existing nanoparticle drug carriers. A typical drug-release profile of nanoparticles suggests that the nanoparticles would immediately release drug into the bloodstream upon intravenous administration and thus only a portion of drugs reach the tumors, causing non-targeted drug release, low drug efficiency, toxicity to healthy tissues and less drug being available to cancer [Liu et al., 2001]. The initial burst release is caused by large surface area of nanoparticles and poorly entrapped drugs, or drugs adsorbed onto the outside of the particles. It has been proposed that the burst release could be minimized by creating chemical bonding of the hydrophobic polymers with the drugs, such as poly(lactic-co-glycolic acid) (PLGA) with a terminal free carboxylic acid group conjugating with DOX [Yoo et al., 2000] or covalently grafting DOX to core-forming poly(aspartic acid) [Kataoka et al., 2001]. Such nanoparticles, however, showed low or completely no anticancer activity [Yokoyama et al., 1998], because chemically bound DOX is not released due to the absence of hydrolysable link between the drug and the polymer chains of the core.
A second issue is their slow drug release. After the initial burst release, the drug release from the nanoparticles became very slow. Cancer cells have many forms of over-expressed drug resistance. If the drug influx into the cancer cell is lower than the capacity of drug removal by the P-gp pumps and the drug metabolism and detoxication etc by cancer cell's other forms of drug resistance, the drug cannot build up a concentration higher than the cell-killing threshold concentration for effective killing. The cores of existing nanoparticles are made of solid polymers and the drugs have to diffuse from the core to the outside and thus the drug release is inevitably slow.
A third issue of the nanoparticles is their slow cellular uptake by cancer cells. Because cancer cells have the P-gp pumps located in the membrane that can effectively transport the drugs out of the cell while they are in the membrane, drugs released in cell interstitium cannot effectively enter the cell plasma through the membrane. In contrast, nanoparticles in the cell release drugs directly in the tumor cell plasma and thus circumvent the P-gp pumps. Therefore, drug release inside cancer cells is preferable, but calls for efficient cellular uptake of the nanoparticles. For the PEG-coated nanoparticles, the PEG layer is used to minimize the nanoparticle interaction with RES cells to evade the clearance by RES for a long blood circulation time, but it also substantially slows down the cancer cellular uptake of the nanoparticles by the same mechanism—the steric repulsion of the PEG chains [Klibanov et al., 1990; Torchilin et al., 1992; Vittaz et al., 1996; De. Jaeghere et al., 2000]. For instance, the majority of localized PEG-coated vesicles were found not to interact with target cancer cells [Yuan et al., 1994]. As a result, PEG-coated nanoparticles may just passively accumulate in cancer interstitium and release drugs there. Receptor-mediated endocytosis by installing ligand moieties on the nanoparticle surfaces has been used to enhance the cellular uptake, e.g. transferrin receptor—transferrin [Ogris et al., 1999; Dash et al., 2000], folate receptor—folic acid [Leamon et al., 1999; Leamon and Low, 2001; Lu and Low, 2002; Kennedy et al., 2003].
Nevertheless, nanoparticles functionalized with folic acid only may not be sufficiently effective to overcome the cancer drug resistance. The steric repulsion of the nanoparticle's PEG outer layer may prevent the folic acid on the nanoparticles from finding and binding the folate receptors. Thus, the internalization of nanoparticles only via FR-mediated endocytosis (FR-Endoc) may not be fast enough to build up a cytoplasmic drug concentration exceeding the capacity of the cell's drug resistance. Hence, FR-Endoc needs to be accelerated. Furthermore, not all cancers express folate receptors. Even in FR-positive tumors, there are cells having low or no FR expression because of the cell heterogeneity. These cells cannot effectively take up the folic acid-functionalized nanoparticles and thus survive the treatment, causing relapse.
As a result of premature burst release by existing nanoparticles, slow drug release rate and low cellular uptake rate, the drug influx into the cancer cell by existing nanoparticles is still lower than the drug efflux and destruction by cell's drug resistance. The drug concentration in cancer cells is thus still lower than the cell-king threshold concentration and cannot effectively induce cell death.
Folate receptors are over-expressed on various types of cancer cells, and mediate endocytosis (FR-Endo) of folic acid-conjugated carriers. Carriers with cationic charges are easily adsorbed onto negatively charged cell membranes and enter the cell via adsorptive endocytosis (AD-Endo). In this invention, folic acid and cationic-charges dually functionalized, lysosomal pH-responsive drug releasing nanoparticles (DFNp) for fast cytoplasmic drug delivery is invented. The dual functions enable the nanoparticles to be efficiently internalized via combined endocytosis mechanisms (FR-Endo, AD-Endo, adsorption-promoted folate-receptor mediated endocytosis (AD-FR-Endo)). The nanoparticle cores are composed of polymers that are soluble at lysosomal pH (˜5) and rupture the lysosomal membrane, and thereby can rapidly release drugs into the cytoplasm. Tumor targeting of the nanoparticles is achieved by passive accumulation through permeable blood capillaries, and active targeting to acidic interstitium and folate receptors. It is believed that these nanoparticles can deliver a large amount of drugs to the cytoplasm overcoming drug resistance for high therapeutic efficacy.